Titanium and titanium alloys have been popularly used in many medical applications due to their light weight, excellent mechanical performance and corrosion resistance. Examples for use of commercially pure titanium (c.p. Ti) include a dental implant, crown and bridge, denture framework, pacemaker case, heart valve cage and reconstruction devices, etc. Nevertheless, due to its relatively low strength, c.p. Ti may not be used for high load-bearing applications.
The most widely-used titanium alloy for load-bearing applications is Ti-6Al-4V alloy (the work-horse titanium alloy). With a much higher strength than c.p. Ti, Ti-6Al-4V alloy has been widely used in a variety of stress-bearing orthopedic applications, such as hip prosthesis and artificial knee joint. Moreover, the lower elastic modulus allows the titanium alloy to more closely approximate the stiffness of bone for use in orthopedic devices compared to alternative stainless steel and cobalt-chrome alloys in orthopedic implants. Thus, devices formed from the titanium alloy produce less bone stress shielding and consequently interfere less with bone viability.
One major potential problem with Ti-6Al-4V alloy as being used as an implant material is its less biocompatible Al and V elements. Studies indicated that release of Al and/or V ions from Ti-6Al-4V implant might cause long-term health problems (Rao et al. 1996, Yumoto et al. 1992, Walker et al. 1989, McLachlan et al. 1983). Its poor wear resistance could further accelerate the release of these harmful ions (Wang 1996, McKellop and RoKstlund 1990, Rieu 1992).
Another problem with c.p. Ti and Ti-6Al-4V alloy is their relatively high elastic modulus. Although their elastic modulus (about 110 GPa) is much lower than the popularly-used 316L stainless steel and Co—Cr—Mo alloy (200-210 GPa), the moduli of c.p. Ti and Ti-6Al-4V alloy are still much higher than that of the natural bone (for example, only about 20 GPa or so for typical cortical bone). The large difference in modulus between natural bone and implant is the primary cause for the well-recognized “stress-shielding effect.”
According to Wolff's law (bone's response to strain) and bone remodeling principles, the ability of a prosthetic restoration/implant construct to transfer appropriate stresses to the surrounding bone can help maintain integrity of the bone. There has been, and still is, a concern about the high elastic modulus of metallic implants compared to bone. Stress shielding phenomenon, more often observed in cementless hip, knee prostheses and spinal implants, can potentially lead to bone resorption and eventual failure of the arthroplasty (Sumner and Galante 1992, Engh and Bobyn 1988).
Both strain gauge analysis (Lewis et al. 1984) and finite element analysis (Koeneman et al. 1991) demonstrated that lower modulus femoral hip implant components result in stresses and strains that are closer to those of intact femur, and lower modulus hip prosthesis better simulates the natural femur in distributing stress to adjacent bone tissue (Cheal 1992, Prendergast and Taylor 1990). Canine and sheep implantation studies revealed significantly reduced bone resorption in the animals with low modulus hip implants (Bobyn et al. 1992). Bobyn et al. (1990, 1992) also showed that the bone loss commonly experienced by hip prosthesis patients may be reduced by using a prosthesis with lower modulus.
It is generally accepted that reduction in Young's modulus value of an implant may improve stress redistribution to the adjacent bone tissues, reduce stress shielding and eventually prolong device lifetime. Metallic implant materials with higher strength/modulus ratios are more favorable due to a combined effect of high strength and reduced stress-shielding risk.
It is known that reduction in Young's modulus value of an implant can reduce stress shielding and prolong device lifetime, and that a metallic implant material with a higher strength/modulus ratio is favorable due to a combined effect of high strength and reduced stress-shielding risk. Nevertheless, from the viewpoint of alloy design, simultaneously increasing the alloy strength and increasing the alloy modulus has always been a great challenge. The strength and modulus of alloys are almost always increased, or decreased, at the same time.
A series of β and near-β phase Ti alloys with better biocompatibility and lower moduli (than Ti-6Al-4V) have recently been developed. Nevertheless, these alloys usually need to contain large amounts of such β-promoting elements as Ta, Nb and W. For example, about 50 wt % and 35 wt % of Ta and Nb, respectively, are needed to form a β-phase binary Ti—Ta alloy and Ti—Nb alloy. Addition of large amounts of such heavy weight, high cost and high melting temperature elements increases the density (Low density is one inherent advantage of Ti and Ti alloys), manufacturing cost, and difficulties in processing.
More recently an Al and V-free, high strength, low modulus α″ phase Ti—Mo based alloy system (typically Ti-7.5Mo) has been developed in the present inventors' laboratory, which demonstrates mechanical properties superior to most existing implantable Ti alloys and a great potential for use as orthopedic or dental implant material.
Biocompatibility of this α″ type Ti-7.5Mo alloy was confirmed through cytotoxicity test and animal implantation study. The cell activity of this alloy is similar to that of Al2O3 (control). Animal study indicates that, after 6 weeks of implantation, new bone formation is readily observed at alloy surface. It is interesting to note that, after 26 weeks, the amounts of new bone growth onto the surface of Ti-7.5Mo implants at similar implantation site are dramatically larger than that of Ti-6Al-4V implant, indicating a much faster healing process.
U.S. Pat. No. 6,726,787 B2 provides the process for making such a biocompatible, low modulus, high strength titanium alloy, which comprises preparing a titanium alloy having a composition consisting essentially of at least one isomorphous beta stabilizing element selected from the group consisting of Mo, Nb, Ta and W; and the balance Ti, wherein said composition has a Mo equivalent value from about 6 to about 9. The key process for obtaining the low modulus, high strength titanium alloys is that the alloys must undergo a fast cooling process at a cooling rate greater than 10° C. per second, preferably greater than 20° C. per second from a temperature higher than 800° C. Said Mo equivalent value, [Mo]eq, is represented by the following equation, [Mo]eq=[Mo]+0.28[Nb]+0.22[Ta]+0.44[W], wherein [Mo], [Nb], [Ta] and [W] are percentages of Mo, Nb, Ta and W, respectively, based on the weight of the composition.
Nevertheless, alloys with a non-cubic (non-symmetrical) orthorhombic crystal structure α″ phase are generally difficult to be cold-worked. The poor cold-workability largely limits the applications of the materials. Titanium alloys with an α″ phase primarily include Ti—Mo based, Ti—Nb based, Ti—Ta based and Ti—W based alloys.